The optimal matching of an artificial intraocular lens requires knowledge of the optical conditions in the patient's eye, in particular the distances between the cornea, crystalline lens and retina.
After this determination of position was originally carried out by means of ultrasound, a device operating optically and without making contact has been introduced in the form of the IOL Master of Carl Zeiss. The functional principle is based in this case on the so called time domain optical coherence domain reflectometry, a short coherence interferometry method such as is described, for example, in WO 00/33729, the content of which is incorporated by reference herein. The main component is a Michelson interferometer that enables the detection of interference of light scattered back by the cornea, lens and retina. The use of a short coherence light source means that it is possible for always only short wave chains to interfere with one another, and this determines the measuring accuracy. So that axial movements of the patient do not falsify the measurement result, the so called dual beam method is applied in which the light scattered back by the cornea serves as reference.
Since the measuring range that must be more than 43 mm for an eye (typical eye lengths vary between approximately 20 and 32 mm, extreme ones between 14 and 40 mm, the mean refractive index being approximately 1.36), must be traversed mechanically by the reference mirror in the case of Michelson interferometer, a measurement usually lasts a few seconds in which the patient is, for example, not allowed to blink since the eyelid movement would render the measurement impossible.
Efforts to accelerate the rate of adjustment of the reference path, for example, by rotating prisms such as EP 1 391 781, have not been successful, since the sensitivity is not sufficient to achieve the required measuring accuracy.
In DE 43 09 056 describes another measurement method based on short coherence, in the case of which light from a broadband light source is shone into the sample, and the light scattered back from various depths is analyzed spectrally. The depth information is obtained from a Fourier transformation of the detected signal. This method is denoted as spectral domain OCDR (SD ODCR) or, because of the Fourier transformation used, also as Fourier domain OCDR (FD OCDR). This category also includes the swept source OCDR (SS OCDR), which is described in the article entitled “High-speed optical frequency-domain imaging” by S. H. Yun et al., Optics Express 2003, page 2953, and in which the light source is tuned spectrally, and the signal received by the detector likewise includes the depth information after the Fourier transformation. As already shown in U.S. Pat. No. 5,321,501 for time domain OCT (TD OCT), the imaging required to implement optical coherence tomography (OCT) is implemented by Galvo scanners that deflect the measurement beam laterally over the sample.
Along the lines of the terminology introduced in the case of the ultrasound measuring device, the one-dimensional (axial) measurement in the case of OCDR along the light axis is generally denoted as an A-scan in general, and therefore also below. Likewise along the lines of the ultrasound terminology, the two-dimensional measurement with the aid of a lateral component in the case of OCT is also denoted as a B-scan.
A first attempt to apply SS OCDR in optical biometry was described in F. Lexer, C. K. Hitzenberger, A. F. Fercher and M. Kulhavy “Wavelength-tuning interferometry of intraocular distances”, Appl. Optics 36 (1997) pages 6548-6553. This solution showed that it is possible in principle to measure the intraocular distances in the eye, although the measuring accuracy was much too inaccurate at 0.82 mm.
An improvement to this solution was disclosed in C. K. Hitzenberger, M. Kulhavy, F. Lexer, A. Baumgartner “In-vivo intraocular ranging by wavelength tuning interferometry”, SPIE [3251-6] 1998. Here, a resolution of 0.15 mm was reached, but it still does not correspond to the requirements. The measuring accuracy for the eye length must, however, be smaller than 30 μm in order to limit the residual errors of the determined IOL refraction to 1/10 diopters.
In particular, the OCDR and OCT methods on moving samples such as, for example, the human eye have the problem that the sample can move during the measurement and this, as discussed in S. H. Yun et al. (2004), OPTICS EXPRESS 2977, can greatly reduce the signals and falsify them. The usual approaches to eliminating the problem are the extremely complicated tracking methods in which the movement of the sample is detected and the measurement beam position is tracked.
Such approaches to the compensation of typical movements of a few hundred micrometers per second are described, for example, in Hammer et al. (2005), Journal of Biomedical Optics 10(2), 024038, and in US 2006/105903. It is disadvantageous of such approaches that, despite the large technical outlay, the finite latency time of such systems always results in certain tracking errors, particularly for very fast eye movements such as saccades.